Magnetic resonance imaging (MRI), nuclear magnetic resonance (NMR), and other magnetic resonance (MR) apparatus, systems, and approaches continue to become more sophisticated, powerful, precise, and complicated. MR involves the transmission of carefully controlled RF energy in the presence of carefully controlled magnetic fields to produce NMR in a material exposed to the RF energy.
Clinical MRI experiences constraints in spatial resolution due to hard physical and physiological limits. Increasing the strength of the magnetic fields used in MRI to, for example, 7T improves spatial resolution but at the expense of reduced contrast. As the magnetic field is strengthened, higher frequencies are needed for the RF to produce NMR because of the Larmor relationship:ω=γB0 
where:                ω is the precession frequency        γ is the gyromagnetic ratio, and        B0 is the magnetic field strength.        
The higher frequencies used with the higher magnetic field strength apparatus yield reduced effectiveness of contrast agents used in MR. For example, a contrast agent that is used to produce a first change in T1 at a lower frequency and field strength may produce a second, lower change in T1 at a higher frequency and field strength. T1 refers to spin-lattice relaxation and T2 refers to spin-spin relaxation.
Due to the physical and physiological limits, conventional 1.5T or 3T human scanners have typically been limited to a resolution of approximately 2×2×2 mm3. However, some targets to be evaluated using MRI (e.g., cancer cells, tumors, proteins) may be significantly smaller than 2×2×2 mm3. For example, some tumor cells may be as small as 10 microns.
Conventionally, even though an MR signal may have been acquired from a tumor or cancer cell that was less than the voxel size used in MR acquisition and reconstruction, it has been difficult, if even possible at all, to distinguish voxels that include small targets (e.g., cancer cells) from voxels that do not include small targets due to the masking effect of non-specific uptake.
The non-specific uptake masking problem arises because conventional contrast agents have a large non-specific enhancement component. The non-specific enhancement issue arises even in specific targeted molecular imaging based contrast agents that only recognize (e.g., bind to) a specific marker. The non-specific enhancement may stem, for example, from vascular or other structural changes in tissue that cause non-specific uptake. The non-specific uptake masking issue is particularly problematic when attempting to image targets (e.g., tumors) smaller than a voxel.
FIGS. 1A and 1B show the baseline MR signal intensities SIG1 and SIG2 associated with two different voxels 110 and 120 that are experiencing NMR before a molecular imaging agent that affects an MR parameter (e.g., T1) has been applied. FIG. 1A illustrates a voxel 110 that includes a small tumor 100 while FIG. 1B illustrates a voxel 120 that is tumor free. Voxels 110 and 120 show a practically indistinguishable total MR signal intensity (e.g. SIG1 and SIG2 are equal).
FIGS. 1C and 1D show total MR signal intensities SIG3 and SIG4 associated with the two different voxels 110 and 120 experiencing NMR after a molecular imaging agent that affects an MR parameter has been applied. The molecular imaging agent may be, for example, a molecular imaging probe that recognizes tumor 100 and that will be taken up by tumor 100 more than it will be taken up by non-tumor tissue. The heavier shading in tumor 100 represents a higher concentration of the molecular imaging agent. The molecular imaging agent may be conjugated to a contrast agent (e.g., Gadolinium (Gd)). SIG3 is a function of baseline signal intensity (e.g., SIG1, SIG2) plus signal intensity due to non-specific uptake plus signal intensity due to specific uptake. SIG4 is a function of baseline signal intensity plus signal intensity due to non-specific uptake. Like SIG1 and SIG2 are difficult, if even possible to distinguish, so too SIG3 and SIG4 are practically indistinguishable. SIG3 and SIG4 are indistinguishable due, at least in part, to the dilution of the MR signal associated with specific uptake in tumor 100 by the MR signal associated with non-specific uptake throughout the voxel 110. Even though tumor 100 shows a large specific uptake of the molecular imaging agent, the signal associated with the specific uptake of the molecular imaging agent in tumor 100 is overwhelmed by the signal associated with non-specific uptake of the molecular imaging agent in the rest of the sample. The non-specific uptake in voxel 110 is so similar to the non-specific uptake in voxel 120, and the total MR signal is so dominated by the non-specific uptake, that MR signal due to specific uptake in tumor 100 is masked making SIG3 and SIG4 practically indistinguishable. Since SIG3 and SIG4 are indistinguishable, no specific characterization (e.g., diagnosis) of material in FIG. 1C (e.g., tumor 100) can be made, even though there was significant specific uptake in the tumor 100.